Imaging devices, such as an x-ray imager, have been used for diagnostic and treatment purposes. One type of x-ray imager is a diagnostic imager configured to operate with a diagnostic radiation source. Another type of x-ray imager is a high DQE detector that is configured for use with a treatment radiation source. An x-ray imager may also be configured for use with both diagnostic radiation beam and treatment radiation beam.
Creating a high DQE detector for portal imaging presents a significant technical challenge. One approach uses thick pixilated scintillator arrays that are coupled to an electronic portal imaging device (EPID). Incoming x-ray photons deposit energy into the scintillators which then produce optical photons via luminescence. These optical photons, which originate with random polarizations and direction vectors after the luminescence events, are transported throughout the scintillator during which time they can be reflected, refracted and scattered. Eventually, many photons will cross the boundary between the scintillator and the photodiode array to be absorbed by the EPID's photodiodes and converted into electrical current for readout and digitization. Despite the promise of the technology, performance may be inadequate and a significant manufacturing cost lies in the process of cutting the crystalline scintillators into parallelepipeds and gluing reflective septa between them in order to reduce optical cross talk.
Also, in some cases, an x-ray imager (e.g., a diagnostic x-ray imager or a portal imager) may comprise a scintillator coupled to a photodiode array. X-ray photons deposit energy into the scintillator thereby producing optical photons with random direction and polarization vectors. A percentage of these optical photons will cross the scintillator-photodiode boundary and deposit energy. The photodiodes convert optical photons into electron-hole pairs. After a sufficient amount of charge is collected, signals are read out and digitized to form an image. To achieve a sufficiently high spatial resolution, optical blurring is desired to be minimized. This implies that the photodiode signals associated with a given x-ray photon should be localized in close lateral proximity to where that x-ray photon interacted with the scintillator. A common means of achieving this goal is through the use of pixelated geometries that confine optical photons using reflective septa. Unfortunately, this approach suffers from high manufacturing costs and may not be practical for incorporating into large-area imagers. As similarly discussed, the process of cutting the crystalline scintillators (e.g. CsI, CdWO4, BGO) into parallelepipeds, gluing reflective septa between them, and then assembling the pixels into a complete array, may be very expensive. Another disadvantage of the pixelated geometry is the loss of fill factor (and associated quantum efficiency) due to the finite thickness of the septa.
Also, current amorphous silicon based flat panel imagers for megavoltage radiation suffers from very low x-ray conversion efficiency. Only about 1.3% of the x-ray photons contribute to an image. In other words, more than 98% of the imaging dose gets lost and will not contribute to the image formation. Approaches that utilize thicker scintillator are either very expensive because the scintillator has to be pixelated or has to exhibit very high imaging performance due to added blurring.